During typical surgical procedures, patient bleeding is controlled not only to avoid excess blood loss, but also for such purposes as assuring that the vision of the performing surgeon is not obstructed or the operation otherwise disrupted. In "open" surgery, bleeding traditionally has been controlled through the use of ties, clamps, blotting, or the like. Further, commencing in the 1920's, a technology emerged referred to as "electrosurgery" in which tissue is cut and blood coagulated with a select application of electrical current. The latter effect is referred to as hemostasis.
Today, electrosurgery is one of the more widely used hemostatic surgical modalities for cutting tissue and carrying out coagulation. Electrosurgical instrumentation falls into one of two categories, monopolar devices and bipoloar devices. Generally, surgeons are trained in the use of both monopolar and bipolar electrosurgical techniques, and essentially all operating rooms will be found equipped with the somewhat ubiquitous instrumentality for performing electrosurgery, the electrosurgical generator.
A somewhat pioneer monopolar electrosurgical device was developed by William T. Bovie. This early device described, for example, in U.S. Pat. No. 1,813,902, issued on Jul. 14, 1931 entitled "Electrosurgical Apparatus" has met with acceptance over the years within the surgical community to the extent that current versions are referred to as the "Bovie." Current such devices typically consist of a handle having a first or "active" electrode extending from one end. The other end of the handle is electrically coupled to an electrosurgical generator which provides a high frequency electric current in either an A.C. cutting mode or a pulsed coagulating mode. A remote control switch is attached to the generator and commonly extends to a foot switch located in proximity to the operating theater. During an operation, a second or "return" electrode, having a much larger surface area than the active electrode, will be positioned in contact with the skin of the patient. To remove tissue, the surgeon brings the active electrode in close proximity with the tissue and activates the foot controlled switch. Electrical current then arcs from the distal portion of the active electrode and flows through tissue to the larger return electrode. In a cutting mode, electrical arcing and corresponding current flow results in a highly intense but localized heating which causes cell destruction and tissue severance. Generally, the device is then switched to a pulsed, higher voltage input to perform in a coagulating mode.
For the bipoloar modality, no return electrode attached to the patient is used. Instead, a second electrode is closely positioned adjacent to the first electrode, with both electrodes being attached to a handle. As with monopolar devices, this handle is electrically coupled to an electrosurgical generator. When this generator is switch activated, electrical current arcs from the end of the first electrode to that of the second electrode. In turn, tissue disposed between the electrodes is cut and blood coagulated. In practice, several electrodes may be employed, and depending on the relative size or locality of the electrodes, one or more electrodes may be active.
Both the monopolar and bipolar devices have inherent shortcomings. For example, since monopolar devices use relatively large radio frequency (RF) currents, the patient may be subject to unwanted risks of RF bums occurring inadvertently at locations where the patient is grounded. Additionally, the path of current flow through tissue often presents the risk of tissue necrosis (i.e., tissue and organ cell death) along errant current flow paths. Equally, tissue necrosis may occur at excessive depths located at the point of application near the active electrode. In any event, a considerable portion of the RF current necessarily is dissipated through the patient, a condition of risk.
By contrast, the bipolar modality overcomes some of the more undesirable characteristics of monopolar devices in that excessive necrosis is reduced, and current is not passed extensively through the body. Since current arcs between adjacent electrodes, blood vessels are readily cauterized. Bipolar devices, however, generally exhibit poor cutting ability. Additionally, it is difficult to accurately locate the arc between the two electrodes with respect to tissue under resection.
Perhaps, as a consequence of the inherent disadvantages associated with monopolar and bipolar modalities, investigators have considered resistively-heated surgical blades or cutting implements to provide a capability for simultaneously cutting tissue and coagulating blood. In their early form, these instruments consisted of a surgical wire or blade which was electrically connected to a power supply. As current passed through the device, the blade was resistively heated to a high temperature. Thereafter, the surgical wire or blade was used to incise tissue but with the added advantage that the generated and localized heat served to cauterize at the incision.
One problem associated with these early thermally based surgical devices is that the wire or blade will rapidly cool upon contact with tissue. As the blade cools, it becomes less and less effective for providing hemostasis. Additionally, as the blade cools below a temperature threshold, tissue tends to stick to it, resulting in obstruction of the cutting edge. If additional power is supplied to accommodate for the cooling effect, overheating may occur in some regions of the blade. Such overheating may be accompanied by unwanted tissue burning or blade destruction. Steel blades exhibit these problems in particular (especially where heating is a function of I.sup.2 R and resistance increases with temperature).
One solution to these problematic heating phenomena has been to provide a surgical blade which has a cutting edge or tissue engaging region with a self-regulating temperature characteristic. Self-regulation (also known as auto-regulation) involves maintaining the cutting edge of the surgical blade within an elevated, preselected temperature range. An approach for attaining self-regulation is has been to employ a ferromagnetic material in constructing the end or heating element of the surgical instrument. When AC current, or more particularly RF current, is passed through such ferromagnetic material, the current density tends to concentrate near its outer surface. This current density attenuates exponentially as distance into the material from the surface increases, a phenomenum known as the "skin effect."
The depth of the skin effect (i.e., the distance of penetrating current density into the ferromagnetic material) is defined as the depth at which current is reduced to approximately 37% of its surface value. This depth may be represented mathematically as follows: EQU skin depth =(5.times.10.sup.3).sqroot..rho./.mu..function.
where skin depth is measured in centimeters, .rho. is electrical resistivity in ohmcentimeters, .mu. is electrical relative magnetic permeability for the ferromagnetic material, and .function. is frequency of the applied alternating electric potential.
In ferromagnetic materials (such as iron, nickel, cobalt, and respective alloys), adjacent atoms and molecules couple their magnetic moments together in rigid parallelism (an interaction known as exchange coupling) in spite of the randomizing tendency of the thermal motion of atoms. If the temperature of such material is raised above a "Curie" temperature, specific for each ferromagnetic material, the noted exchange coupling suddenly disappears. As a result, these materials exhibit large changes in relative permeability as the temperature of the ferromagnetic material transitions through the Curie temperature. As seen from the above-mathematical equation, since the relative permeability is known to change in response to the temperature of the material, the associated skin depth also will change. This relationship between skin depth and temperature enables ferromagnetic material based instruments to achieve auto-regulation.
The heating elements of surgical devices can be constructed from ferromagnetic material which is selected to have a Curie temperature at or near the auto-regulation temperature desired for a particular surgical application. As RF current passes through the ferromagnetic material, the heating element will resistively heat to approximately the Curie temperature. Once the cutting edge contacts tissue, both it and the area surrounding it will cool to a level below the Curie temperature. In response to this Curie transition, skin depth will decrease which, in turn, results in an increased resistance of the cooled region (the resistance being a function of the ferromagnetic material's resistivity multiplied by length and divided by area). A corresponding increase in power supplied will accompany this increase in resistance. The temperature then will tend to again increase due to resistive heating toward the Curie temperature. Thus, auto-regulation of the surgical blade or wire around the Curie temperature is achieved.
By way of example, the heating element of a surgical device may be formed from an iron-nickel alloy having a Curie temperature of about 450.degree. C. While the heating element is above the Curie temperature, the relative permeability (shown in the above equation) remains near one. However, when the cutting portion of the heating element contacts tissue, the temperature will correspondingly drop and the relative permeability will rise, for example, to 100 or even 1000. The associated skin depth, in turn, decreases by more than 10 to 1. Correspondingly, an increase in power is supplied to the heating element, which increases resistive heating --thus achieving auto-regulation.
A disadvantage associated with resistively heated devices, including those which employ ferromagnetic heating elements, is concerned with a lack of sufficient localization of heat at the actual cutting edge. In this regard, the entire heating element including the support for its cutting edge is heated toward a Curie temperature suited for cutting. This poses a risk that the support portion of the heating element may contact tissue or organs not selected for incision. Additionally, since a larger portion of the heating component is heated, the time period required for the cutting region to cool down to safe levels posing no threat of burning can be quite significant. This time period, for example, may be five seconds or more, an interval generally considered unacceptable for surgical procedures. Further, during a laparoscopic or endoscopic procedure, the view of the surgeon is confined to a camera-generated two dimensional image at a monitor, such as a TV screen. The heated element, however, may be moved out of the camera's limited field of view during the several seconds which are required for cool-down, thereby endangering tissue that may come in contact therewith.
Another disadvantage associated with resistively heated devices is concerned with a perceived requirement that they must be powered by a specially designed power supply, typically incorporating a constant current source. Such electrical drive systems generally are configured to be unique to the properties of a particular heating element and are not of a universal nature, such that they would be usable with different resistively heated tools. For example, some devices require not only a constant current source but a complex control which senses a voltage decrease as the cutting edge approaches Curie temperature. Thereafter, the control system turns the current source off for a predetermined time to reduce the temperature of the cutting edge. Of course, such added equipment requirements pose budgetary concerns to health care institutions.
Essentially all hospitals, health care centers, clinics, and the like which have facilities for performing surgical procedures are equipped with an ample number of conventional electrosurgical generators. These generators are designed to operate in monopolar and bipolar modes to, in turn, drive conventional instruments. In this regard, the generators commonly have two output sets which accommodate monopolar devices and bipolar devices. The generators, however, are not compatible with resistively heated surgical devices.
It therefore would be desirable to provide a resistively-heated surgical instrument where only a confined region of the heating element reaches the elevated temperatures required for cutting and coagulation of tissue and where the heating element rapidly cools to temperatures atraumatic to tissue and organs. Any connective member attached to the heating element for providing support thereto should remain at temperatures below a threshold which would otherwise produce unwanted thermal damage to surrounding tissue.
It also would be desirable to provide a low-cost resistively-heated surgical instrument which can be operated using conventional electrosurgical generators already in wide use in operating facilities.
In addition to the above, it further would be desirable to provide a resistively-heated surgical instrument that overcomes patient risks heretofore associated with monopolar electrosurgical devices and which improves the cutting ability heretofore associated with bipolar electrosurgical devices.
It also would be desirable to offer a resistively-heated surgical instrument with replaceable working tips to allow the surgeon to select a specific working end configuration and size particularly suited to the tissue cutting and coagulating requirements at a given stage of a surgical procedure.
It also would be desirable to provide a resistively-heated surgical instrument which has the ability to precisely control the location and amount of thermal energy delivered to tissue.